Patent Application
20220395614
A portion of the disclosure of this patent document contains material that is subject to copyright protection. The copyright owner has no objection to the facsimile reproduction by anyone of the patent document or the patent disclosure, as it appears in the Patent and Trademark Office patent files or records, but otherwise reserves all copyright rights whatsoever.
A biomimetic, polymeric composite biomaterial designed as a heart valve leaflet substitute that can be used for heart valve repair and/or to fabricate a new-generation of durable heart valve prosthesis. In some embodiments, the polymeric composite biomaterial is in the form of a patch.
Valve replacement in adults and children has inherent problems associated with anticoagulation (mechanical valves) or durability (bioprosthetic heart valves), which leads to the failure of the prosthesis and increases the probability for reoperation and the accompanying risk. Thus, valve repair is always the preferred approach, compared to replacement.
Valve repairs frequently require the use of cardiovascular patches to perform leaflet augmentation or extension. Current available patches used for valve repair, such as bovine pericardium, porcine intestinal submucosa extracellular matrix, expanded polytetrafluoroethylene, fresh autologous pericardium and glutaraldehyde-treated autologous pericardium, all have intrinsic limitations and drawbacks that affect their long-term durability and mechanical performance, leading to structural degeneration (SD) of the patch and of the repaired valve leaflet.
Baird et al. used PhotoFix® patches in young patients for valve repair and showed cases of degeneration, calcification and inflammation. (Baird, C. W. et al. Photo-Oxidized Bovine Pericardium in Congenital Cardiac Surgery: Single-Centre Experience. Interact CardioVasc Thorac Surg 2017, 24 (2), 240-244). Hofmann et al. reported a high rate of mechanical failure of CorMatrix® patches for aortic valve repair leading to valve insufficiency. CardioCel® patch was also associated with a significant risk of patch failure and the need for reoperation in large series. (Hofmann, M. et al., Congenital Aortic Valve Repair Using CorMatrix®: A Histologic Evaluation. Xenotransplantation 2017, 24 (6), e12341). Pavy et al. summarized that the discrepancy between the mechanical property (the elasticity) of the patch and the native tissue was the result to cause severe aortic stenosis in infants and lead to patch failure. (Pavy, C. et al. Initial 2-Year Results of CardioCel® Patch Implantation in Children. Interactive CardioVascular and Thoracic Surgery 2018, 26 (3), 448-453), It is also in line with Tremblay's conclusion, which reported that the difference in the mechanical properties of aortic tissues and prosthetic material was a main factor to contribute the unwanted hemodynamic effects leading to patch failure. (Tremblay, D. et al. Comparison of Mechanical Properties of Materials Used in Aortic Arch Reconstruction. The Annals of Thoracic Surgery 2009, 88 (5), 1484-1491).
To overcome these drawbacks, implantable patches or grafts with native-like structures and tunable mechanical performance close to the ones of the native valve leaflets were attempted. In this regard, a polyvinyl alcohol (PVA)-bacterial cellulose (BC)-based hydrogel was designed to mimic the mechanical properties of the native valve leaflet. But, the degradable nature of PVA over time, poor design of the patch and lack of data on the durability of the composite hindered its application as a stable cardiac patch. Another composite fabrication involved the combination of poly(ethylene glycol) (PEG) hydrogel and polycaprolactone (PCL) fiber for heart valve tissue engineering, but this composite material demonstrated an anisotropic behavior on the unicycle tensile test only and had a linear stress-strain behavior that was different from the non-linear behavior of native leaflets. This may cause valvular interstitial cells (VICs) to experience greater stresses, impact VIC activation and extracellular matrix (ECM) remodeling, leading to calcification.
Masoumi et al. attempted a tri-layered scaffold designed to mimic structural and anisotropic mechanical characteristics of the native leaflet. (Masoumi, N. et al., A. Tri-Layered Elastomeric Scaffolds for Engineering Heart Valve Leaflets. Biomaterials 2014, 35 (27), 7774-7785). But this biodegradable patch degraded at a fast rate with a loss of mechanical strength from 3.02 MPa to 1.63 MPa in 4 weeks. In comparison, native valve aortic and pulmonary leaflets keep a stable modulus of, 3.84 and 2.55 MPa, respectively over time. Despite aiming at replicating the architecture of native leaflets, Masoumi's patch is not a mechanically stable option for clinical use. Thus, there remains a need for a stable, functional and biomimetic patch that overcomes the drawbacks of previous patches, including a clinical need for a new type of patch that can achieve a better durability after implantation in patients
In one aspect, a stable biomimetic polymeric biomaterial is provided. The biomaterial includes at least two layers including a Fibrosa-mimic (“F-mimic”) layer, a Spongiosa-mimic (“S-mimic”) layer, and a Ventricularis-mimic (“V-mimic”) layer. In some embodiments, the F and V layers are anisotropic and the S layer is a shock absorbing layer. In some embodiments, the F-mimic layer and the V-mimic layer are made of polycarbonate polyurethane (PCU) film, enhanced with aligned, electrospun polycaprolactone (PCL) fibers, and the S-mimic layer is made of PCU foam. In some embodiments, the stable biomimetic polymeric biomaterial includes two to five layers. The stable biomimetic polymeric biomaterial may be devoid of animal-derived tissue, thus, in some embodiments it has no animal-derived tissue. The biomaterial may be used to make a patch, such as for treating a heart defect, or a prosthetic heart valve.
In another aspect, a polymeric, biomimetic customized biomaterial patch (“BCP”) that replicates the structure-function driven architecture of native valve leaflets is provided and described herein. In one embodiment, the BCP replicates the three-layer architecture and the anisotropic mechanical properties of a native leaflet. Thus, in one embodiment, the BCP comprises a composite body including three polymeric layers. In this regard, the layers include a Fibrosa-mimic (“F-mimic”) layer; a Spongiosa-mimic (“S-mimic”) layer; and a Ventricularis-mimic (“V-mimic”) layer. In some embodiments, the F and V layers are anisotropic and the S layer is a shock absorbing layer. In some embodiments, the F-mimic layer and the V-mimic layer are made of polycarbonate polyurethane (PCU) film, enhanced with aligned, electrospun polycaprolactone (PCL) fiber mesh, and the S-mimic layer is made of PCU foam. In some embodiments, the biomimetic patch is entirely polymeric, i.e., lacks animal-derived tissue. The tri-layered patch can be modified and tuned to achieve the specific mechanical requirement, as well as have a low antigenicity and lower risk for structural valve degeneration.
The BCP described herein, as compared to three commercial patches, exhibits an anisotropic mechanical behavior and mechanical stiffness (6.20±1.83 MPa and 1.80±0.21 MPa in circumferential and radial directions, respectively), which is more similar to the native aortic valve leaflets than any currently available commercial patches. The BCPs also exhibits greater durability and greater biocompatibility. In vivo rat subcutaneous tests also confirmed the BCP exhibits mechanical biostability and superior resistance to inflammation and calcification, compared to the commercial patches. Thus, the BCP embodied herein provide a new clinical-grade biomaterial patch useful for heart valve extension, augmentation or replacement in children and adults.
In yet another aspect, a novel polymeric valved device, such as an implantable prosthetic heart valve is provided. The implantable prosthetic heart valve comprises the biomaterial described herein. The implantable prosthetic heart valve may be an aortic valve, mitral valve, or tricuspid valve.
In yet a further aspect, methods for repairing a heart defect with the BCP and methods for delivering an implantable heart valve to a subject in need thereof is described and embodied herein.
Various embodiments of the present disclosure can be further explained with reference to the attached drawings, wherein like structures are referred to by like numerals throughout the several views. The drawings shown are not necessarily to scale, with emphasis instead generally being placed upon illustrating the principles of the present disclosure. Therefore, specific structural and functional details disclosed herein are not to be interpreted as limiting, but merely as a representative basis for teaching one skilled in the art to variously employ one or more illustrative embodiments.
Various detailed embodiments of the present disclosure, taken in conjunction with the accompanying figures, are disclosed herein; however, it is to be understood that the disclosed embodiments are merely illustrative. In addition, each of the examples given in connection with the various embodiments of the present disclosure is intended to be illustrative, and not restrictive.
Throughout the specification, the following terms take the meanings explicitly associated herein, unless the context clearly dictates otherwise. The phrases “in one embodiment” and “in some embodiments” as used herein do not necessarily refer to the same embodiment(s), though it may. Furthermore, the phrases “in another embodiment” and “in some other embodiments” as used herein do not necessarily refer to a different embodiment, although it may. Thus, as described below, various embodiments may be readily combined, without departing from the scope or spirit of the present disclosure.
In addition, the term “based on” is not exclusive and allows for being based on additional factors not described, unless the context clearly dictates otherwise. In addition, throughout the specification, the meaning of “a,” “an,” and “the” include plural references. The meaning of “in” includes “in” and “on.”
As used herein, the terms “and” and “or” may be used interchangeably to refer to a set of items in both the conjunctive and disjunctive in order to encompass the full description of combinations and alternatives of the items. By way of example, a set of items may be listed with the disjunctive “or”, or with the conjunction “and.” In either case, the set is to be interpreted as meaning each of the items singularly as alternatives, as well as any combination of the listed items.
In one aspect, a biomimetic polymeric biomaterial is provided. The biomimetic polymeric biomaterial is useful as a heart valve leaflet substitute and/or to fabricate a prosthesis. In one embodiment, the polymeric biomaterial is used to make a biomimetic customized biomaterial patch or BCP. In other embodiments, the biomimetic polymeric biomaterial is used to fabricate a polymeric valve prosthetic device. Thus, the biomimetic polymeric biomaterial may be used for treating a subject in need of heart valve repair and/or heart valve replacement.
Generally, the BCP comprises a body having a multi-layered polymeric composite biomaterial. For example but not limitation, the multi-layered polymeric composite biomaterial may include two to five layers. In one embodiment, the biomaterial is a tri-layered polymer composite. In this embodiment, the BCP is designed to mimic the architecture, i.e., three distinct tissue layers that compose the valve leaflets, and the mechanical properties of native leaflet tissue.
Referring to
Referring to
Referring to
It has been found that the biomaterial comprising the S-mimic foam layer and a plurality of polypropylene fibers embedded in the foam structure to form a composite biomaterial offers mechanical properties substantially the same as native leaflet tissue, as shown in Table 1A below. As shown, the biomaterial in some embodiments exhibits a tensile modulus in a C/H direction of about 8 to about 16 MPa.
Referring to
Exemplary Materials and Method for fabricating embodiment of
The BCP 100 was prepared by a combination of three native-tissue mimicking layers, respectively named the Fibrosa-mimic 101 (F-mimic) layer, the Spongiosa-mimic (S-mimic) layer 102 and the Ventricularis-mimic (V-mimic) layer 103. The F-mimic layer and V-mimic layer were designed as fiber-enhanced layers, composed of aligned PCL fibers and PCU film in order to replicate the anisotropy of these layers. The S-mimic layer was designed as a PCU foam to replicate the load-bearing mechanical role played by the native spongiosa. The structure of this BCP is shown in
Fabrication of the F-mimic 101 and the V-mimic 102 fiber enhanced layers. Using a 15% PCL solution prepared in a mixed solvent (Chloroform:Methanol=3:1), PCL fibers were produced by electrospinning with the following parameters: a flow of 1 ml/hour, a voltage of 20 kV voltage and a distance of 15 cm between the nozzle and drum collector. The solution was spun towards a rotating collector at a rate of 1600 rpm to collect the aligned fibers. The fibers were allowed to dry overnight in a chemical hood for solvent evaporation before the following fabrication and characterization. The collected, aligned PCL fibers were embedded in solution-casted PCU film. The 15% PCU solution was casted by a doctor-blade coater through a 500 μm gap to control the film thickness. The fiber-solution composite was cured overnight in a chemical hood to evaporate the solvent and form the fiber-enhanced layers.
Fabrication of the S-mimic layer 103. To produce the S-mimic layer, the 15% PCU solution was casted by a doctor-blade coater to create a film with a fixed thickness of 1500 μm. Subsequently, the film was immersed in deionized water for 24 hours. Then, the solvent-exchanged PCU film was frozen under −80° C. Lyophilization was conducted on the frozen PCU film at 0.1 mBar, −40° C. for 72 hours and turned into a porous layer to work as the S-mimic layer.
Fabrication of the BCP 100. The F-mimic layer was casted to form the fiber-enhanced layer. After 1 hour drying in the hood, the S-mimic layer was put over the casted composite and dried with the fiber-enhanced layer together in the chemical hood. Then this two-layer composite was put over the V-mimic layer to fabricate the BCPs.
Morphology Characterization. To characterize the PCL aligned fibers, each mimic layer and the BCPs, the specimens were sputter coated with gold/platinum and imaged with a Zeiss Sigma VP scanning electron microscope (SEM) at an accelerating voltage of 3 kV. SEM images were used for the visual inspection of fiber's orientation, mimic layer and BCPs' inner structures and their surface quality.
Tensile mechanical testing. The mechanical properties were measured with an Instron 5848 mechanical tester with a 50 N load cell at a strain rate of 10% s-1. The specimens were cut as 5 mm×20 mm stripes (for non-tissue samples) or 3 mm×10 mm ones (for the native tissue samples) in two different directions, horizontally/circumferentially (H or C direction) and vertically/radially (V or R direction), shown in
where
are engineering stress and engineering strain. l0, w0, and T0 are the dimension (length, width and thickness) of the specimen, Δl is the change of elongation in length, and ΔF is the change of the force. Then the average curves and the tensile modulus at the strain of 15% or 40% were used to compare the mechanical performance in different directions and to assess anisotropy. Then the average curves and the tensile modulus at the strain of 15% or 40% were used to compare the mechanical performance in different directions and to assess anisotropy.
Flexural mechanical testing. Flexural properties of the commercial patches, leaflet tissues and BCPs were tested via the bulge tests. All samples were pre-cut as the circular planar specimens using fine dissectors. Thickness was evaluated by averaging three measurements taken at specimen's center with a digital caliper. The diameter of the caliper's contact plate was 10 mm, which was larger than the circular test area with a diameter of 6 mm; thus the specimens were assumed to be uniform in thickness. The specimens were speckled with black India ink to allow for DIC deformation tracking. The specimens were then glued between two plates with holes of 6 mm diameter (
The specimens were inflated by a custom-made displacement-driven syringe injection of PBS into the custom-made pressurization chamber. The pressure was monitored by a pressure transducer with 0-8 kPa range. The loading regimen was programmed using LabView (V2020, National Instruments, Austin, Tex.) and displayed in
The deforming specimen surface was imaged by two stereoscopically arranged cameras with 20 mm focus lengths at an aperture of f/4. The optical axes of the cameras were positioned 35 cm above the chamber and fixed with a total angle of 12°. This configuration had a depth of field in front over 1.5 cm, sufficient to capture the deformation of the specimen between 0.2-7.2 kPa. Images were collected during testing at a rate of 10 Hz by VicSnap 2009 and correlated by Vic3D (V8, Correlated Solution, Inc. Columbia, S.C., USA).
To calculate the pressure and displacement resultants, the measured pressure and displacement were tared by the baselines, resulting in relatively zero stress and strain at the reference state. The method of calculating the flexural modulus Eflex is provided below. The sample in this test was modeled as a circular thin plate with edges fully fixed. The pressure was evenly distributed on the bottom surface of the sample. The governing equation and boundary conditions of this case could be expressed in cylindrical coordinates (r, ø, z) as
where w is the displacement of z direction (defined as the out-of-plane direction) at a point of the thin plate, R is the radius of the plate, and ΔP is the pressure exerted. D is the flexural rigidity defined as
T0 is the thickness of the specimen. The solution to this equation is
At the center point (r=0), the flexural modulus could be expressed as
where ΔW is the change of the displacement in z-direction. Here, all the materials were assumed to be incompressible, so the Poisson's ratios were all set as 0.5.
Suture retention testing. The suture retention capabilities of the three commercial patches, PCU films and the BCPs were tested following the steps described in Pensalfini et al.'s work, using Instron 5848 tensile machine (
and compared among all the samples.
Biostability testing. Specimens of the commercial patches, PCU films/foams and BCPs were pre-cut as 5 mm×30 mm and submerged into 2 mL vials filled with an in vitro solution of 20% hydrogen peroxide (H2O2)/0.1M cobalt chloride (CoCl2). The in vitro solution was refreshed twice a week, and all testing were done at 37° C. After a period of 5, 10, 14, 15, 20, 24, and 30 days, the specimens were removed, rinsed thoroughly in deionized-water, dried in the hood, then cut into two parts (5 mm×25 mm and 5 mm×5 mm). The former was tested via the tensile mechanical testing and the tensile modulus at strain=15% was calculated. The latter was analyzed by SEM to inspect the surface quality.
Biocompatibility testing. Bovine Serum Albumin (BSA) static protein-adsorption experiments. For static protein-adsorption tests, 1 mg mL-1 BSA solution was prepared in PBS (pH 7.4). Commercial patches and PCU films were cut into specimens (50 mm×10 mm) and immersed in 10 mL 1 mg mL-1 BSA solution in a test tube. BSA adsorption was conducted under vibration at 37° C. for 3 hours to allow for adsorption equilibrium. Then the specimens were rinsed with PBS, the remaining proteins adsorbed on the surfaces were removed with a 1 wt % aqueous solution of sodium dodecylsulfate (SDS), similar to the work done by Song et al. The experiments were performed with five measurements for each specimen. BSA content was measured using a NanoDrop™ spectrophotometer at a wavelength of 280 nm and then the amount of adsorbed BSA on specimens was calculated.
Calcium-ion (Ca2+) adhesion experiments. The Ca2+ adhesion experiments were performed in a metastable calcium phosphate (MCP) solution. The purpose of using this MCP solution is to obtain calcium-phosphate compounds which can precipitate out from the solution and deposit on the tested specimens, in order to test the samples' calcification resistance in in vitro studies. Similar experiments were performed as reported earlier. In brief, 3.87 millimole (mM) CaCl2), 2.32 mM K2HPO4 and 0.05M Tris buffer were solved in 1000 ml of de-ionized water, to yield a Ca/PO4 ratio of 1.67.
This solution is more physiologically representative of hydroxyapatite, with a Ca/PO4 ratio of 1.67, which is the most common form of calcium minerals in the vascular calcification process. Commercial patches and PCU films were cut into specimens (5 mm×30 mm) and immersed in 2 mL MCP solution individually. This experiment was conducted under vibration at 37° C. and solution was changed every 48 hours to ensure an adequate ion concentration. The specimens were removed after 16 days and rinsed with water to remove excess solution and loosely attached deposits. The specimens were dried in the vacuum oven at 70° C. overnight, accurately weighed, and hydrolyzed in 2 mL of 2 N HCl for 24 hours at 50° C. The calcium concentration was determined from HCl hydrolysate, using calcium colorimetric assay.
Rat subcutaneous implant model. In accordance with NIH guidelines for the care and use of laboratory animals (NIH Publication #85-23 Rev. 1985), all animal protocols were approved by the Institutional Animal Care and Use Committee (IACUC) of Columbia University (Protocol #AC-AABD5614).
Eighteen specimens (diameter=8 mm) of PCU film (n=6), Gore-Tex® patch (n=6) and CardioCel® patch (n=6) were implanted in the subcutaneous position of three rats. Following induction of anesthesia, fur clipping, and standard sterile prepping and draping, six subcutaneous pockets were created on the dorsal surface of each rat. One specimen was implanted into each pocket, after which all wounds were re-approximated with surgical clips. The rats were sacrificed at 8 weeks with an overdose of isoflurane (Euthenase).
Histology. The implanted specimen was retrieved while still contained in host tissue, fixed in 10% neutral buffered formalin and processed using paraffin-embedding techniques. Slides were stained with Hematoxylin and Eosin and Alizarin Red stains. In each specimen, both the patch and the surrounding host tissue were evaluated.
Calcium Content & mechanical test. Samples were analyzed for calcium content using calcium colorimetric assay as described in Calcium-ion adhesion experiments described above. Briefly, the specimen disks were removed from host tissue, fixed in formalin and solvent-exchanged in DI-water. Following with the lyophilization, the net weight of the specimen disks were acquired. After hydrolyzing in nitric acid, the calcium content was quantitated. Results are reported as microgram calcium per milligram dry specimen weight. The PCU disks can be separated from the host tissue after lyophilization. This specimen's mechanical performance was also evaluated as described in section 2.4 and its tensile modulus at the strain=15% was recorded, to compare with the control, unplanted sample.
Statistical Analysis. Statistical analyses of the tensile mechanical properties, biostability mechanical tests, protein adsorption and calcium adhesion tests were performed using one-way analysis of variance (ANOVA). P values less than 0.05 were considered statistically significant (*P<0.05, **P<0.01, ***P<0.001 and ****P<0.0001). And differences between samples within the groups were evaluated using a student's t-test, or Tukey's multiple comparisons test followed by ANOVA. GraphPad Prism 7 (San Diego, Calif., USA), statistics package, was used to obtain statistical significance for the study above.
Results. Structure and Mechanical Properties. Tensile properties of native tissues and commercial patches. We performed cyclic, uniaxial tensile tests on native leaflets and commercial patches in order to compare the mechanical performance of our BCP with these reference tissues (
Mechanical characteristics of commercial patches (Gore-Tex®, CorMatrix® and CardioCel®) were obtained under the same conditions as the native tissues and are presented in Table 1. Compared to native tissues, the commercial patches are generally much stiffer, with a tensile modulus in the range of 6-120 MPa at strain=15% in C-direction and 23-180 MPa at strain=40% in R-direction. Commercial patches also display a non-anisotropic behavior, with similar tensile modulus at the same strain level in the horizontal (H) and vertical (V) directions (Table. 1,
Referring to
Structure and tensile properties of the mimic layers and the BCPs.
In order to increase the compliant of the overall composite, also to correspond to the Spongiosa layer, PCU foam was made via lyophilization. Compared to the film made from the same concentration PCU solution, the foam exhibited a porous structure (
The biomimetic, customized three-layered composite patch (BCP) was obtained by coating the fiber-enhanced layers on both sides of the S-mimic layer (
Referring to
Flexural properties. Bulge tests were performed to assess the flexural properties of the HAV, commercial patches and the BCP disclosed and embodied herein. The bulge test measured the components of the displacements in a 3D coordinate (
Preconditioning was found to have a negligible effect on the mechanical response.
Anisotropic level.
Referring to
Table 3 summarized the ratio of principal strain and second principal strain in-plane, e1/e2. Among the three commercial patches, the Gore-Tex® patch was the most isotropic one, and CorMatrix® is the most anisotropic. For native tissues, PPV and HAV have obvious anisotropic behaviors. For BCPs, although from the design and the tensile test data they were demonstrated as the anisotropic composites, the average ratio of e1/e2 is just higher than Gore-Tex®. It may be attributed to the similar scale of the tensile modulus in-plane X and Y directions.
Flexural Modulus: Table 3 also summarized the data of thickness and displacement of specimens in the out-of-plane direction. It can be seen that the commercial patches generally possessed higher flexural modulus: Gore-Tex® was the stiffest among those three types of patches (17.58±4.50 MPa) and CardioCel® was the most compliant one (4.52±2.40 MPa). Native tissues, including porcine leaflets and human leaflets, behaved more compliant than commercial patches during the bulge tests. For BCPs, they had a similar flexural modulus range (3.55±2.80 MPa) as HAV (2.70±1.30 MPa), and displayed better compliance than Gore-Tex® and CorMatrix®.
Suture retention: The resistance to tearing of the BCPs, the raw material (PCU film) and the three commercial patches were determined by suture retention strength measurements. The mean suture retention strength (SRS) of Gore-Tex®, CardioCel® and CorMatrix® are 5.35±1.25 N, 8.99±1.77 N and 4.07±1.38 N respectively (Table 4). The SRS of the BCP and the PCU film are in the range of the commercial patches (
A thickness-normalized SRS (TN-SRS) has also been applied to eliminate the effect of sample thickness and needle size. According to the Equation 5, the TN-SRS of Gore-Tex®, CardioCel® and CorMatrix® were 94.76±22.14 N/mm2, 98.26±19.35 N/mm2, and 82.86±28.10 N/mm2 respectively (Table 4). There's no significant difference on TN-SRSs among those three commercial patches. TN-SRSs of the BCPs in two directions were 89.91±13.25 N/mm2, and 79.1±11.1 N/mm2 respectively and were not significantly different from the three commercial patches (
Referring to
Biostability. The biostability of the commercial patches, PCU-based raw film/foam and our BCPs were assessed by an accelerated oxidative degradation test, using a 0.1 M CoCl2/20% H2O2 solution.
Referring again to
Biocompatibility. BSA Protein Adsorption. A BSA protein adsorption test was applied to assess the blood compatibility of the three commercial patches and BCPs.
Ca2+ Adhesion. Table. 5 shows the results of 16-day Ca2+ adhesion tests performed on the PCU film, BCP and commercial patches. It has clearly stated that the PCU film and BCP have a lower Ca2+ deposition compared to Gore-Tex® and CardioCel® patches (
Referring to
In vivo studies. Subcutaneous implantation. A set of schematic illustrations of H&E images from three samples: PCU film, Gore-Tex® and CardioCel® Patch, are presented in
PCU film had no evidence of calcification as indicated by
Subsequently a calcium content assay was conducted and confirmed the histological findings. A significant increase in Ca2+ level was found in Gore-Tex® and CardioCel® samples compared to the PCU film, with p<0.0001 (Table 5. and
Referring back to
Sections of three samples also displayed the distribution of calcification (red color shown as shading in the schematic illustrations) in tissues and the patch samples. No visible calcification appeared in PCU specimen but a high degree of calcification presented in two commercial specimens (131-133)
Utilizing the biostable and biocompatible polymers as the main components, described herein is a polymer-based, tri-layered patch to mimic the three-layer architecture of native leaflets. The in vitro and in vivo assessment of our BCPs covers two main parts: the long-term mechanical and biological performance.
The mechanical assessment utilizes the cyclic uni-axial tensile tests, flexural bulge tests and suture retention tests for characterization. Tensile test offers a more direct and more economical approach to characterize the mechanical properties. Studies on the uniaxial tensile properties of valve leaflets in the literature have stretched the specimens to break, and recorded the ultimate stress (MPa), the strain-to-failure/ultimate strain (%), as well as calculated the elastic modulus (MPa) using the Equation 1. The ultimate stress and stain-to-failure were acquired beyond the physiological level and unveiled the properties which were not fit the working range; and the one-time tensile stretch cannot reflect the performance at steady state, especially considering the fact that the initial tensile curve behaves more differently from the rest of cyclic curves of viscoelastic materials due to the Mullins' effect and the preconditioning effects. Tensile modulus, was commonly used to easily quantify an intrinsic elastic property of soft, viscoelastic biomaterials. Note that the stress and strain in the tensile modulus were engineering stress and engineering strain, so the effect of the cross-sectional contraction was not reflected in the tensile modulus.
A 20-time cyclic tensile test for all the samples was conducted. The maximum strain was set as 15% in H/C direction and 40% in V/R direction, corresponding to the physiological level from systole to diastole. The averaged, post-conditioning curves was picked to do the calculation of tensile modulus to eliminate the influence of the Mullin's effect and the preconditioning effects. In order to compare BCP with reference tissues and commercial patches, the tensile modulus was calculated at the strain of 15% and 40% for the H/C-direction and V/R-direction, respectively.
The averaged tensile curves and modulus data display that HAV is stiffer than PAV, and PPV is stiffer than PAV. It was also found that the anisotropic behavior and matched mechanical properties at the specific strain range were hardly achieved in commercial patches. Most of them are either too stiff (except for CardioCel® in H direction) or isotropic, compared to the HAV. They are far from the satisfactory material to match the native tissue, from the mechanical view.
BCP, thus, was designed and fabricated using solution casting, lyophilization and electrospinning to replicate the complex, structure-function driven architecture of native leaflets. It was demonstrated that a patch with such structure (
The flexural properties of the BCP, heart valve tissues and commercial patches, were also studied using the bulge tests. Due to the limitation of the pressure transducer and the capacity of the customized syringe pump, the maximum pressure can reach 7.2 kPa (54 mmHg) as a valid, stable level and a frequency of 0.25 Hz allows for specimen inflation. The results from each specimen still demonstrated the various performance on flexural deformation under a quasi-physiological simulation. Three commercial patches displayed a randomly anisotropy performance and higher flexural modulus (4.52-17.58 MPa). Gore-Tex® is made of ePTFE and has no particular design for anisotropic applications. It leads to an isotropic behavior during the tests. While CorMatrix® and CardioCel® are derived from bio tissues and it is reasonable to have some residual fibers in the patch, which provide anisotropy. For the BCP, due to the similar scale of the tensile modulus in-plane X and Y directions, it didn't display an obvious anisotropic performance in-plane. It also emphasizes the significance to decrease the modulus of the BCPs in V/R direction in order to compare with native tissue level. All of HAV, PAV and BCP have a lower flexural modulus between 0.53-3.55 MPa. This performance is also in line with the trend of tensile modulus data shown in Table 1 and 2, especially the one in C/H direction as shown in
Suture retention capability. Punctures and defects are generated during suturing, which may result in mechanical failure through crack propagation. Therefore, the resistance to tear, characterized as SRS and TN-SRS, are essential to evaluate the feasibility of the patches or alternatives. From the results it can be seen the SRS of our BCP and its raw materials (6.25-6.58 N) were in the range of the ones of commercial patches (4.07-8.99 N), which demonstrates that they have a similar capacity of resistance to tearing as the commercial products. It is noted that a number of different suture thread thicknesses and needle types were applied in the clinics, depending on the detailed applications and surgeons' selection. Some geometrical parameters such as the diameter of the suture, the thickness of the graft wall remain unconstrained by the norm. Thus, TN-SRS was also introduced to evaluate the suture retention capability of the products, normalizing this parameter without impact from the product and thread thicknesses. The TN-SRS of BCP has no significant difference from the ones of commercial patches. And BCP also has a higher toughness than most of commercial patches, which emphasizes its durable nature. To sum up, a series of suture retention tests demonstrated that the BCP has a resistance to tearing similar, even better than the commercial patches, no matter from the SRS, TN-SRS or toughness.
The biological assessment of the BCPs and commercial patches includes the biostability and biocompatibility, in vitro and in vivo. As a designed, polymer-based patch, it is expected to be stable in vivo and the mechanical properties do not alter over time. Published papers reported that the degradation of polyurethane-based materials in vitro and in vivo was attributed to several mechanisms including metal ion-induced accelerated oxidative degradation, hydrolytic degradation and enzymatic degradation. It is demonstrated that oxidative degradation was the more dominant mechanism over other degradations. Thus, a 0.1 M CoCl2/20% H2O2 solution was applied in this test to accelerate oxidative degradation of the PCUs. The Co2+ ions have been demonstrated to rapidly decompose hydrogen peroxide via the Haber-Weiss reaction. Degradation results after 24 days in this solution was shown to correlate to 12 months of in vivo implantation. The modulus of the BCP and PCU film/foam displayed no significant change (NS, One-way ANOVA) on mechanical properties in 30 days in this accelerated oxidization solution. It demonstrated that the BCP has a stable performance which was equivalent to 15 months of in vivo implantation. Even so, a slow oxidative degradation sign was found on the outside surface layer. This finding suggests that the biostability of the BCP, although being comparable to the one of FDA-approved Gore-Tex® patches, may be improved down the road through a surface modification process targeting the resistance to oxidation.
To evaluate biocompatibility, protein adsorption and calcium-ion adhesion are selected to assess BCP and commercial patches' biological performance in vivo. Protein adsorption is a significant factor to determine the thrombogenicity of an implanted graft. When blood gets in contact with the graft's surface, protein adsorption occurs first, then leads to more plugs aggregation, eventually provokes the generation of the fibrin network and thrombus formation. Thus, our BCP should aim at reducing their potential for protein adsorption and cut the path of forming thrombin. Bovine serum albumin has a structure similar to human serum albumin (HSA) and the HSA has the highest concentration in human plasma. A BSA protein adsorption test was performed to characterize the blood compatibility of the surfaces of our BCPs and the commercial patches. And the BCP exhibited a low level of protein adsorption compared to three commercial patches. It may be attributed to its smooth PCU film surface without holes or sites, which avoids the plugs deposition and formation.
On the other hand, it is significant to evaluate the resistance to calcification when developing any biomaterial since calcification is the leading reason of failure of bioprosthetic heart valves and grafts. It is a complex phenomenon influenced by a series of mechanical and biochemical factors. It also limits the durability of synthetic polymer materials used in heart valve devices and blood contact application in general. In vitro Ca2+ adhesion tests using a MCP solution to mimics the hydroxyapatite level were performed. A 16-day test exhibits that the BCP and its main component PCU film have a lower level of Ca2+ ion accumulation compared to commercial patches. And this trend is also in line with the findings from in vivo subcutaneous tests (
An in vivo rat subcutaneous implantation has been conducted to verify the biostability and biocompatibility of PCU film (the main composition of BCPs) and two commercial patches. The former exhibited a stable performance and little/no cell or tissue infuse or grow within the patch. It also exhibited a little-to-no calcification level, better than commercial patches. No obvious mechanical properties degradation after the tests and the tissue generated around the PCU patch were organized and no-calcification. It is a good sign to highlight the feasibility to apply the PCU-based BCP in vivo and expect the positive outcomes.
Compared to three commercial patches, this BCP demonstrated an anisotropic mechanical behavior and mechanical stiffness (6.20±1.83 MPa and 1.80±0.21 MPa in circumferential and radial directions, respectively), which was much closer to the native aortic valve leaflets than any currently available commercial patches. What's more, our BCPs also showed an excellent durability in an in vitro accelerated oxidization solution and displayed an excellent biocompatibility with an in vitro lower protein adsorption level and a lower calcium adhesion level. In vivo rat subcutaneous tests confirmed its main composition, PCU's mechanical biostability and superior resistance to inflammation and calcification, compared to the commercial patches.
The native-like performance of the BCP avoids patch failure and degeneration, which are related to the inadequate mechanical properties. It is biostable, and does not rely on uncontrolled polymer degradation and tissue formation. The biomimetic patch also exhibits a low protein adsorption and low Ca2+ adhesion, avoiding a high risk of thrombogenicity and calcification. In some embodiments, fiber meshes can be fabricated by various biocompatible polymers, to optimize the anisotropic mechanical performance.
In some embodiments, the biostability and biocompatibility is optimized through adding the surface layer on the current version, for example, Parylene C can be evenly coated on the patch through chemical vapor deposition.
The biomimetic patch is scalable. For example, at a lab scale, this version of the patch is processed through solution casting, electrospinning and lyophilization. A multiple technology combination provides flexible tuning methods for optimization. At an industrial scale, this tri-layer composite can be fabricated via a non-expensive and scalable multi-layer co-extrusion technology. This green, non-solvent involved method provides a better reproducibility and lower costs of production. It provides a feasible path to commercialize this polymeric patch to improve the durability and quality of the valve repair, and decrease the number of reoperations and complications. In some embodiments, the surface morphology is further processed to create the “corrugations” structure to mimic the native leaflet's surface morphology. This structure plays an important role and accounts for the native collagen fiber's mechanical behavior during valve closing.
Example: A polycarbonate urethane-based material with aligned polycaprolactone fibers to enhance the anisotropic properties are disclosed. Solution casting, electrospinning and lyophilization were used to mimic the native leaflet's architecture. Compared to current commercial materials, this BMM exhibited an anisotropic behavior and a mechanical performance much closer to the native aortic leaflets. The material exhibited biostability in an accelerated oxidization solution equivalent to 15 months of implantation. It also displayed better resistance to protein adsorption and calcification in vitro and in vivo. This material is shown to have long-term durability for surgical valve repair or replacement.
Materials: Carbothane™ AC-4075A, Polycarbonated-based polyurethane (PCU) (Lubrizol, Wilmington, Mass.) was dissolved in dimethylacetamide (DMAC) (Acros Organics, Fair Lawn, N.J.). Polycaprolactone (PCL, Mw=80,000; Sigma-Aldrich, St. Louis, Mo.) was used to create fibers and dissolved with a mix of chloroform (Sigma-Aldrich, St. Louis, Mo.) and methanol (Fisher Scientific, Hampton, N.H.) with a 3:1 molar ratio. Three commercially available patches were selected for comparison: Gore-Tex® (W. L. Gore and Associates, Flagstaff, Ariz., USA), CorMatrix® (Cardiovascular, Inc, Atlanta, Ga., USA) and CardioCel® (Admedus, Toowong, Queensland, Australia). The CryoValve® aortic human valve (CryoLife Inc., Kennesaw, Ga., USA) was used as the control sample, after being dissected and kept intact in PBS.
Fabrication of the biomimetic multilayered material (BMM). Fibrosa-mimic layer and Ventricularis-mimic layer fabrication: Using a 15% PCL solution prepared in a mixed solvent, PCL fibers were produced by electrospinning with the following parameters: a flow rate of 1 ml/hour, a voltage of 20 kV voltage and a distance of 15 cm between the nozzle and drum collector. The solution was spun towards a rotating collector at a rate of 1600 rpm to collect the aligned fibers. The fibers were dried overnight in a chemical hood 192 for solvent evaporation. The collected, aligned PCL fibers were embedded in a solution-casted PCU film. The 15% PCU solution was casted by a doctor-blade coater through a 500 μm gap to control the film thickness. The fiber-solution composites, fibrosa-mimic (F-mimic) layer and ventricularis-mimic (V-mimic) layer, were cured overnight in a chemical hood to evaporate the solvent and form the fiber-enhanced layers.
Spongiosa-mimic layer fabrication: 15% PCU solution was casted by a doctor blade coater to create a film with a fixed thickness of 1500 μm. Subsequently, the film was immersed in deionized water for 24 hours in order to replace the solvent with water. The film was frozen at −80 and lyophilized at 0.1 mBar and −40 for 72 hours, leading to the formation of a porous structure (or foam)
BMM fabrication: The F-mimic layer was used as the bottom layer of the BMM. It was fabricated first as described above. After 1 hour drying in the hood, the spongiosa-mimic (S-mimic) layer was placed over the half-cured composite and fully cured with this F-mimic layer overnight. Then this two-layer composite was stacked on top of the V-mimic layer to fabricate the three-layered BMMs using the same strategy.
Morphology characterization. The specimens (PCL aligned fibers, each mimic layer and the BMMs) were sputter coated with gold/platinum and imaged with a Zeiss Sigma VP scanning electron microscope (SEM) at an accelerating voltage of 3 kV. SEM images were used to assess the fibers' orientation, mimic layers and the BMMs' structures and surface morphology.
Tensile mechanical tests. Mechanical tests were performed using an Instron 5848 mechanical tester with a 50 N load cell at a strain rate of 10% s-1. The specimens were cut as 5 mm×20 mm stripes (for non-tissue samples) or 3 mm×10 mm ones (for the native tissue samples) in two different directions, circumferentially (C-direction) and radially (R direction). The thickness was measured at three different points with a digital caliper (Mitutoyo America Corp, Aurora, Ill., USA) and the values were averaged. Four to six specimens for each sample were repeatedly stretched for 20 cycles, either to a maximal strain of 15% in the C-direction or to a maximal strain of 40% in the R-direction. Missirlis and Chong, Brewer et al, Thubrikar et al. and Li et al. have all reported in vivo AV leaflet strains of physiological level to be approximately 10-15% and 30-40% in the circumferential and radial directions respectively. After the first 5 preconditioning cycles, the subsequent 15 cycles of stress-strain curves were recorded and averaged and the tensile modulus E were calculated as the Equation 1, discussed herein.
Flexural mechanical tests. Samples for the flexural mechanical tests were cut as planar specimens with enough area to fully cover the test hole (diameter=6 mm). Thickness was evaluated by averaging three measurements taken at specimen's center with a digital caliper. The specimens were speckled with black India ink to allow for digital image correction (DIC) tracking deformation and glued between two plates with holes of 6 mm diameter (
Specimens were inflated by a custom-made displacement-driven syringe injection of PBS into the custom-made pressurization chamber. The pressure was monitored by a pressure transducer with a range of 0-8 kPa. The loading regimen was programmed using Lab View (V2020, National Instruments, Austin, Tex.). The specimen was brought to a baseline pressure of 0.2 kPa and held for 30 seconds prior to cyclic testing to ensure the 250 specimen was at equilibrium. The specimens were subjected to 30 load-unload cycles at a rate of 3.5 kPa/s from the baseline pressure to a maximum pressure of 7.2 kPa (
The flexural bulge test measured the components of the displacements in a 3D coordinate plane, providing the U, V and W components of the displacement field in X, Y and Z directions. The elastic modulus measured with this flexural bulge test, Eflex, was calculated through the change of the applied pressure (ΔP) and the change of the out-of-plane displacement component (ΔW). The sample in this test was modeled as a circular thin plate with edges fully fixed. The pressure was evenly distributed on the bottom surface of the sample. The governing equation and boundary conditions of this case could be expressed in cylindrical coordinates (r, as in Equation [2] described herein. The solution to this equation is derived as equation [3] described herein. At the center point (=0), Eflex. is expressed as equation [4]. Where ΔW is the change of the displacement in z direction. Here, all the materials were assumed to be incompressible, so the Poisson's ratios were all set as 0.5.
Suture retention tests. The suture retention tests were conducted using an Instron 5848 tensile machine. Prolene 5-0 suture was inserted 2 mm from the end of the 10×15 mm specimen and through the specimen to form a half loop. The suture was pulled at a rate of 50 mm/min crosshead speed. Five specimens were tested in each group. The force (N) required to pull the suture through and/or cause the specimen to fail was recorded as the suture retention strength (SRS). A thickness normalized suture retention strength (TN-SRS, N/mm2) was also applied to eliminate the effect of sample thickness and needle size. TN-SRS is calculated by dividing the suture retention strength by the area of the sample over which the load was applied, Equation [5] as described herein.
Biostability tests. Specimens were pre-cut as 5 mm×30 mm and submerged into 2 mL vials filled with an in vitro solution of 20% hydrogen peroxide (H2O2)/0.1M cobalt chloride (CoCl2). The in vitro solution was refreshed twice a week, and all tests were done at 37° C. After a period of 5, 10, 14, 15, 20, 24, and 30 days, the specimens were removed, rinsed thoroughly in deionized water, dried in the hood, then cut into two parts (5 mm×25 mm and 5 mm×5 mm). The former was tested via the tensile tests and the modulus at strain=15% was calculated. The latter was analyzed by SEM to inspect the surface quality.
Biocompatibility tests. Bovine Serum Albumin (BSA) static protein-adsorption experiments. For static protein-adsorption tests, 1 mg mL-1 BSA solution was prepared in PBS (pH 7.4). BMMs and commercial patches were cut into specimens (50 mm×10 mm) and immersed in 10 mL 1 mg mL-1 BSA solution in a test tube. BSA adsorption was conducted under vibration at 37° C. for 3 hours to allow for adsorption equilibrium. Then the specimens were rinsed with PBS, and the remaining proteins adsorbed on the surfaces were removed with a 1 wt % aqueous solution of sodium dodecylsulfate (SDS), as described by Song et al. The experiments were performed with five measurements for each specimen. BSA content was measured using a NanoDrop™ spectrophotometer at a wavelength of 280 nm, and then the amount of adsorbed BSA on specimens was calculated.
Calcium deposition experiments. The calcium deposition experiments were performed in a metastable calcium phosphate (MCP) solution. The MCP solution has been previously described in detail. In brief, 3.87 millimole (mM) CaCl2), 2.32 mM K2HPO4 and 0.05M Tris buffer were solved in 1000 ml of de-ionized water, to yield a Ca/PO4 ratio of 1.67. This solution is more physiologically representative of hydroxyapatite, which is the most common form of calcium minerals in the vascular calcification process. BMM, PCU film and commercial patches were cut into specimens (5 mm×30 mm) and immersed in 2 mL MCP solution individually. This experiment was conducted under vibration at 37° C., and the solution was changed every 48 hours to ensure an adequate ion concentration. The specimens were removed after 16 days and rinsed with water to remove excess solution and loosely attached deposits. The specimens were dried in the vacuum oven at 70° C. overnight, weighed and hydrolyzed in 2 mL of 2 N HCl for 24 hours at 50° C. The calcium concentration was determined from HCl hydrolysate, using a calcium colorimetric assay.
Rat subcutaneous implant model. In accordance with NIH guidelines for the care and use of laboratory animals (NIH Publication #85-23 Rev. 1985), all animal protocols were approved by the Institutional Animal Care and Use Committee (IACUC) of Columbia University (Protocol #AC-AABD5614).
Eighteen specimens (diameter 333=8 mm) of PCU film (n=6), Gore-Tex® patch (n=6) and CardioCel® patch (n=6) were implanted in the subcutaneous position of three rats. Following induction of anesthesia, fur clipping, standard sterile prepping and draping, six subcutaneous pockets were created on the dorsal surface of each rat. One specimen was implanted into each pocket, after which all wounds were re-approximated with surgical clips. The rats were sacrificed at 8 weeks with an overdose of isoflurane (Euthenase).
Histology. The implanted specimen was retrieved while still contained in host tissue, fixed in 10% neutral buffered formalin and processed using paraffin-embedding techniques. Slides were stained with Hematoxylin and Eosin and Alizarin Red stains. In each specimen, both the patch and the surrounding host tissue were evaluated.
Calcium content & mechanical test. Samples were analyzed for calcium content using calcium colorimetric assay as described above regarding biocompatibility tests. Briefly, the specimen disks were removed from host tissue, fixed in formalin and solvent-exchanged in DI-water. Following the lyophilization, the net weight of the specimen disks was acquired. After hydrolyzing in nitric acid, the calcium content was quantified (microgram calcium per milligram dry specimen weight). The PCU disks were separated from the host tissue after lyophilization. This specimen's mechanical performance was also evaluated as described above regarding tensile mechanical testing, and its tensile modulus at the strain=15% was recorded, to compare with the control, unimplanted sample.
Statistical Analysis. For studies including various groups of samples, like the tensile property studies, the suture retention tests, the biocompatibility studies (protein adsorption and calcium deposition), etc., two-sided t-tests for parametric data with Welch's correction were conducted and used for analysis. The biostability studies, which included the same group of BMM samples, were analyzed using the one-way ANOVA followed by Tukey's post-hoc tests (GraphPad Prism 7, San Diego, Calif., USA). Results are showed as means±standard deviation. P values less than 0.05 were considered statistically significant (*P<0.05, **P<0.01, ***P<0.001 and ****P<0.0001).
The results of the testing are discussed herein. Structure. The BMM was designed as a tri-layer polymeric structure that was specifically developed to mimic the tri-layer anatomy of the native valve (
Tensile properties. Cyclic, uniaxial tensile tests were performed to assess the tensile properties of BMM and its component layers (Table 6).
The aligned PCL fibers exhibit a highly anisotropic performance (35.74±9.81 MPa vs. 1.63±0.38 MPa), compared to the random PCL fibers electrospun from the same solution (7.37±0.30 MPa). Due to the incorporation of the aligned PCL fibers, the fiber-enhanced layers also demonstrate an anisotropic behavior, stronger along the fiber-aligned direction (green solid curve) and similar performance to pure PCU film along the fiber perpendicular direction (green dash curve), shown in
Native tissues and commercial patches were also tested under the same conditions to compare with BMM. For native leaflets, the average stress-strain loading curves exhibit a residue deformation 398 and then an increase in the slope of the stress-strain curves which is attributed to the deformation and stretch of fiber networks. This increase is accentuated in the C-direction compared to the R-direction because of the existence of oriented collagen fibers (
Flexural properties. Bulge tests were performed to assess the flexural properties of the HAV, commercial patches and our BMM. Table 7 summarized the data of thickness and displacement of specimens in the out-of-plane direction. The commercial patches generally possessed higher Eflex: Gore-Tex® was the stiffest among the three commercial patches (16.73±4.28 MPa) and CardioCel® was the most compliant (4.25±2.26 MPa). HAV was more compliant than commercial patches during the bulge tests. BMMs had a similar Eflex range (2.99±2.43 MPa) as HAV (2.54±1.22 MPa), and displayed better compliance than commercial patches. Commercial products have a generally stiffer performance than native tissues and BMMs, whether from tensile tests in two directions or from flexural bulge test (
Suture retention of samples. The resistance to tearing of the BMM and its main component, PCU film, compared to the three commercial patches was determined by suture retention strength (SRS) measurements. The mean SRS of the BMM and the PCU film were 6.58±0.97 N and 6.25±0.88 N respectively. There was no significant difference on SRS of the BMMs in two directions, reflecting a uniform resistance to tearing. The mean SRS of Gore-Tex®, CardioCel® and CorMatrix® were 5.35±1.25 N, 8.99±1.77 N and 4.07±1.38 N respectively (
Biostability. The biostability of the BMMs, PCU film/foam, and three commercial patches were assessed via measuring the degradability of samples in the accelerated oxidative solution. The polymer-based samples (BMM, PCU film/foam and Gore-Tex®) remained stable throughout the 30 days in the accelerated oxidization solution (
Biocompatibility. Bovine Serum Albumin (BSA) adsorption. A BSA protein adsorption test was applied to assess the blood compatibility of the artificial material surface.
Ca2+ deposition To study the material's susceptibility to calcification, the in vitro deposition of Ca2+ ions on BMM, Gore-Tex®, and CardioCel® samples was evaluated in a MCP solution, as previously described.
In vivo studies: subcutaneous implantation. The rat subcutaneous implant model was used for screening cellular infiltration, inflammation and calcification resistance in vivo.
Sections of the three samples also displayed signs of calcification (red and auburn color) in materials and their surrounding tissues (
A calcium content assay was subsequently conducted to confirm the histological findings. A significant increase in Ca2+ level was found in Gore-Tex® and CardioCel® samples compared to the PCU film (
Heart valve leaflets have a highly organized architecture with three specific layers. The fibrosa and ventricularis consist of circumferentially oriented collagen fibers and radially oriented elastin sheets, which constitute their primary load-bearing properties. The spongiosa is inherently soft and compliant with a much lower stiffness. It acts as a cushion, absorbing the load resulting in minimal stress. In this present work, we designed and fabricated a biomimetic, multilayered material to replicate the architecture of those specific layers: The fiber-enhanced PCU films are used as F-mimic layer and V-mimic layers to provide the appropriate mechanical strength and anisotropic properties. A PCU-foam was fabricated via a lyophilization process to create the porous structure from the same polymer solution. It was applied to replicate the load-bearing mechanical role, confer flexibility and tune the overall mechanical properties of BMM. PCU is known as a biostable and biocompatible polymer for heart valve and vascular graft applications. It was found to have superior resistance to degradation under biological conditions when compared with common poly(ether urethane) (PEU) and poly(ether urethane urea) (PEUU). The selection of Carbothane™ AC-4075A as our PCU resin was not only because of its biostable nature, but also due to its mechanical properties in the range of the native tissue (
The mechanical assessment utilized cyclic uni-axial tensile tests, flexural bulge tests, and suture retention tests for characterization. For the tensile test, our averaged stress-strain curves and modulus data displayed that anisotropic behavior and mechanical properties of native HAVs were not achieved by the commercial patches. Compared to the native tissue, three selected commercial patches are either too stiff or isotropic and are therefore far from a satisfactory material to match the native tissue: Gore-Tex® is the expanded polytetrafluoroethylene (e-PTFE) made through a thermal extrusion and it has the most homogeneous performance (e.g. isotropic) among the three commercial patches. It is also the stiffest sample since the carbon atoms in the ePTFE chain are enclosed within a sheath of fluorine atoms. CorMatrix® and CardioCel® are two tissue-derived products: The former is composed of porcine small intestinal submucosa extracellular matrix and the latter is a tissue-engineered ADAPT bovine pericardial patch. Both of them are less stiff due to the tissue nature, and CardioCel® even has a similar tensile modulus to those of HAV in the C-direction. BMM, in comparison, demonstrated a superior, stable performance with valve-mimicking architecture, anisotropic behavior, and stable tensile modulus. The capability of BMM to match the mechanical performance of the native tissue is important to optimize leaflet stresses and decrease tears and perforations. Mismatched properties, especially high stiffness from a rigid material, will lead to fibrosis, inflammation, and loss of elasticity and functionality.
For the flexural properties, bulge tests were first introduced to study the native leaflet tissue and its artificial alternatives in the literature, to the best of our knowledge. The three commercial patches generally displayed either isotropic or uncontrolled and variable anisotropic performance, and they possessed much higher flexural modulus than HAVs. Both BMM and HAV have a lower flexural modulus and this performance is also in line with the trend of tensile modulus data, especially in the C-direction (
Punctures and defects are generated during suturing the artificial materials, leading to mechanical failures through crack propagation. The resistance to tearing is therefore essential to evaluate the feasibility of patches or alternative materials. The SRS and TN-SRS measurements exhibited that the BMMs have a comparable tear resistance to the commercial products. The BMM also displayed a higher toughness (Table 8) than most commercial patches, which emphasizes its durability and capacity to withstand more tear energy than other samples during suturing. A customized heart valve prototype is also fabricated via suturing the BMMs to the 3D-printed valve struts (
The biological assessment of the BMMs and commercial patches included their biostability and biocompatibility, in vitro and in vivo. PCU was selected due to its expected stable and compatible in vivo profile and its stable mechanical properties over time. The degradation of polyurethane-based materials in vitro and in vivo was attributed to several mechanisms, including metal ion-induced accelerated oxidative degradation, hydrolytic degradation and enzymatic degradation. It has been demonstrated that oxidative degradation is the more dominant mechanism over others. Consequently, a 0.1 M CoCl2/20% H2O2 solution was applied to accelerate oxidative degradation of the PCUs. Degradation results after 24 days is shown to correlate to 12 months of in vivo implantation. The modulus of the BMM and PCU film/foam displayed no significant change in mechanical properties for 30 days in this accelerated oxidization solution. It demonstrated that the BMM has a stable performance which is equivalent to 15 months of in vivo implantation. Using polycarbonate macrodiols as the soft segments, the PCU is designed with better hydrolytic stability and anti-oxidization capability than PEU and PEUU. A stable mechanical performance is essential to maintain the mechanical functionality of the valve or patch over time, and to avoid potential failure and repeated reimplantation procedures. These results confirm the biostability of the BMM, which is comparable to the biostable FDA-approved patch (Gore-tex®). However, it is also noted that minor signs of oxidative degradation (
Serum protein adsorption and calcium deposition were examined to evaluate the samples' biocompatibility. Protein adsorption is a significant factor to determine the thrombogenicity of an implanted material. When blood gets in contact with the material's surface, protein adsorption occurs first, which can then provoke the adhesion of platelets and immune cells on the protein layer. Platelets may aggregate continuously and eventually lead to the generation of a non-soluble fibrin network and thrombus formation. An ideal valve leaflet material should have a low protein adsorption profile to limit or cut the path of thrombin formation and potential subsequent thrombogenic reactions. We performed a BSA adsorption test and found that the BMM exhibited a lower level of protein adsorption compared to three commercial patches. Although the difference is not significant, BMM (main composition PCU) displayed a lower surface tension with improved hydrophilicity (
The implantation of any artificial material inevitably provokes a host response. The formation of encapsulated tissue (stained as the pink color in
Compared to three commercial patches, the BMM of the disclosed subject matter demonstrated an anisotropic mechanical behavior and mechanical stiffness which was much closer to the native aortic valve leaflets than the commercial patches. This BMM also showed an excellent durability in an in vitro accelerated oxidization solution and displayed excellent biocompatibility with a lower in vitro protein adsorption level and a lower calcium deposition level. In vivo rat subcutaneous modeling confirmed the mechanical biostability and superior resistance to inflammation and calcification of the main component material, PCU, compared to the commercial patches. This BMM is useful for surgical valve repair and polymeric surgical or transcatheter valve device.
While the disclosure has been illustrated and described in detail in the drawings and foregoing description, such illustration and description are to be considered illustrative or exemplary and not restrictive. The disclosure is not limited to the disclosed embodiments. Variations to the disclosed embodiments and/or implementations can be understood and effected by those skilled in the art in practicing the claimed disclosure, from a study of the drawings, the disclosure and the appended claims.
This application is a Continuation-in-Part of International Application No. PCT/US2020/067002, filed Dec. 24, 2020, which claims priority to U.S. Provisional Application No. 62/953,768 filed Dec. 26, 2019 and U.S. Provisional Application No. 62/976,252 filed Feb. 13, 2020, the disclosures of which are incorporated by reference herein in their entirety. This application claims priority to U.S. Provisional Application No. 63/310,688, filed Feb. 16, 2022, the disclosure of which are incorporated by reference herein in their entirety.
| Number | Date | Country | |
|---|---|---|---|
| 62953768 | Dec 2019 | US | |
| 62976252 | Feb 2020 | US | |
| 63310688 | Feb 2022 | US |
| Number | Date | Country | |
|---|---|---|---|
| Parent | PCT/US2020/067002 | Dec 2020 | US |
| Child | 17849545 | US |
